The invention relates to Magnetic Resonance Imaging (MRI systems, and particularly to the Radio-Frequency (RF) coils used in such systems.
The following references are incorporated herein by reference.
A. U.S. Patent Documents
B. Other References
(1) xe2x80x9cNovel Two Channel Volume Array Design for Angiography of the Head and Neckxe2x80x9d, Reykowski, et al., SMR 2nd Annual Meeting, San Francisco, Calif., p.216 (1994).
(2) xe2x80x9cA Phased Array Coil Optimized for Carotid Artery Imagingxe2x80x9d, Bernstein, et al., ISMRM Seventh Annual Meeting, Philadelphia, Pa., p163 (1999).
Magnetic Resonance Imaging (MRI) utilizes hydrogen nuclear spins of the water molecules in the human body, which are polarized by a strong, uniform, static magnetic field of the magnet (named B0xe2x80x94the main magnetic field in MRI physics). The magnetically polarized nuclear spins generate magnetic moments in the human body. The magnetic moments point in the direction of the main magnetic field in a steady state, and produce no useful information if they arc not disturbed by any excitation.
The generation of the Nuclear Magnetic Resonance (NMR) signal for MRI data acquisition is accomplished by exciting the magnetic moments with a uniform Radio-Frequency (RF) magnetic field (named the B1 field or the excitation field). The B1 field is produced in the imaging region of interest by an RF transmit coil which is driven by a computer-controlled RF transmitter with a power amplifier. During excitation, the nuclear spin system absorbs magnetic energy, and its magnetic moments precess around the direction of the main magnetic field. After excitation, the precessing magnetic moments will go through a process of Free Induction Decay (FID), releasing their absorbed energy and returning to the steady state. During free induction decay, NMR signals are detected by the use of a receive RF coil, which is placed in the vicinity of the excited volume of the human body. The NMR signal is the secondary electrical voltage (or current) in the receive RF coil that has been induced by the precessing magnetic moments of the human tissue. The receive RF coil can be either the transmit coil itself, or an independent receive-only RF coil. The NMR signal is used for producing images by using additional pulsed magnetic gradient fields, which are generated by gradient coils integrated inside the main magnet system. The gradient fields are used to spatially encode the signals and selectively excite a specific volume of the human body. There are usually three sets of gradient coils in a standard MRI system, which generate magnetic fields in the same direction of the main magnetic field, varying linearly in the imaging volume.
In MRI, it is desirable for the excitation and reception to be spatially uniform in the imaging volume for better image uniformity. In a standard MRI system, the best excitation field homogeneity is usually obtained by using a xe2x80x9cwhole-bodyxe2x80x9d volume RF coil for transmission. The xe2x80x9cwhole-bodyxe2x80x9d transmit coil is the largest RF coil in the system. A large coil however, produces lower signal-to-noise ratio (SNR or S/N) if it is also used for reception, mainly because of its greater distance from the signal-generating tissues being imaged. Since a high signal-to-noise ratio is desirable in MRI, special-purpose coils are used for reception to enhance the S/N ratio from the volume of interest.
In practice, a well-designed specialty RF coil should have the following functional properties: high S/N ratio, good uniformity, high unloaded quality factor (Q) of the resonance circuit, and high ratio of the unloaded to loaded Q factors. In addition, the coil device should be mechanically designed to facilitate patient handling and comfort, and to provide a protective barrier between the patient and the RF electronics. Another way to increase the SNR is by quadrature reception. In this method, NMR signals are detected in two orthogonal directions, which are in the transverse plane or perpendicular to the main magnetic field. The two signals are detected by two independent individual coils which cover the same volume of interest. With quadrature reception, the SNR can be increased by up to {square root over (2)} over that of the individual linear coils.
In Magnetic Resonance Imaging (MRI and Magnetic Resonance Angiography (MRA), a carotid RF coil is used to generate high resolution and good SNR images for the carotid arteries with a coverage of about 16 cm. In addition, it is also desirable to use the same carotid coil to cover the arteries for a full field of view (FOV) (about 46 cm for the most of the patient population) from the circle of Willis to the aortic arch with good SNR and image uniformity. For the carotid artery imaging, the performance (i.e., SNR and image resolution) of a carotid coil should be better than that of a neurovascular coil. For the full FOV imaging from the circle of Willis to the aortic arch, the SNR and image uniformity of a carotid coil should be comparable to those of a neurovascular coil. It is also desirable for a carotid coil to be able to do bilateral imaging and unilateral imaging as well.
To cover the blood vessels from the circle of Willis to the aortic arch, a quadrature RF coil was built by Misic, et al. (U.S. Pat. No. 5,517,120). This neurovascular coil utilizes multiple horizontal conductors and end conductors to distribute the current such that two orthogonal magnetic modes (i.e., one horizontal field and one vertical field) are created by the coil to achieve the quadrature detection of the magnetic resonance signal.
A split-top, four channel, birdcage type array coil was also developed by Srinvasan, et al. (U.S. Pat. Nos. 5,664,568; 5,602,479) for head, neck and vascular imaging. This split-top head and neck coil consists of a birdcage head coil (Hayes, U.S. Pat. No. 4,692,705) and two distributed type (flat birdcage type) coils: one for the anterior neck-torso and the other for the posterior neck-torso. The quadrature signal obtained with the head coil is separated into two channels. The anterior and posterior neck-torso coils form the other two channels. The inductive coupling between the neck-torso coils and the head coil is minimized by overlapping the neck-torso coils with the head coil.
The development of array coil technology (Roemer, et al., U.S. Pat. No. 4,825,162) allows one to image a large field-of-view (FOV) while maintaining the SNR characteristic of a small and conformal coil. Using this concept, a two-channel (four linear coils) volume array coil for magnetic resonance angiography of the head and neck was implemented by Reykowski, et al. The first channel is a four bar quadrature head coil consisting of two linear coils. Two Helmholtz type coils form the second channel for covering the neck and chest. The two Helmholtz type coils are arranged such that the magnetic fields generated by them are diagonally oriented and perpendicular to each other (i.e., a quadrature coil pair). The quadrature neck coil is attached to the quadrature head coil. Each of the two Helmholtz type neck coils overlaps with the head coil to minimize the inductive coupling between the head and neck coils, i.e., the neck coils are critically coupled to the head coil, to reduce the noise correlation caused by the cross-talk between the head and the neck coils.
By realizing that the largest part of the carotid arteries lie near the surface of the skin, one can obtain good SNR and high resolution images of the carotid arteries by using smaller size surface coils placing very close the skin rather than using large volume type coils, a 6-element phased array surface coil was developed by Bernstein, et al. dedicated for carotid artery imaging. This carotid coil is flexible. It lies on the patient""s sternum, wraps around the neck and extends to the each side of the head. The carotid coil consists of four 10 cm diameter loop coils, two on each side of the head, for imaging the carotid arteries from the shoulder up (i.e., toward the superior direction or the circle of Willis direction) and two rectangular shape loop coils of size of 4xc3x976 cm and 5xc3x978 cm, respectively, for imaging the arteries from the shoulder down (i.e., toward the inferior direction or the aortic arch direction). Each pair of the two 10 cm diameter loop coils is combined (or multiplexed) into one channel. Therefore, the pair of the two 10 cm diameter loop coils on the right hand side forms one channel to cover the right carotid artery, the pair on the left hand side forms the second channel to cover the left carotid artery and the rectangular loop coils (i.e., the 4xc3x976 cm and 5xc3x978 cm ones) form another two channels to cover the arteries at the chest region. It has been shown that this carotid coil can provide much better SNR and higher resolution for imaging the arteries around the carotid bifurcation than does a commercial neurovascular coil. The total coverage of this carotid coil is about 24 cm.
In R. Srinivasan (U.S. Pat. No. 6,150,816) a three-element mutually decoupled RF coil system, for example, 3 loop coils, was discussed. The two smaller loop coils are overlapped to isolate them from each other. The third large loop coil, which is the same size as the two combined small loop coils, is superimposed on the two small loop coils and physically connected to them. This can be done such that these three loop coils isolate from each other simultaneously. A procedure of iterating testing was developed for optimizing all the three coils of the integrated system. Therefore, the three coils can be turned on simultaneously to receive signals without interfering with each other.
The neurovascular coils (Misic, et al., U.S. Pat. No. 5,517,120 and Srinvasan, et al., U.S. Pat. Nos. 5,664,568 and 5,602,479) and head-neck coil (Reykowski, et al.) are volume type coils. These coils use large coil elements to cover the entire imaging volume (i.e., head and neck). The coil elements are too far away from the carotid arteries to provide good enough SNR and are too large to provide high enough resolution for the carotid artery imaging.
The 6-element phased array surface coil system (Bernstein, et al.) has a limited coverage range of about 24 cm. This coverage is not enough for imaging the arteries from the circle of Willis to the aortic arch for most of the patient population. Therefore, this coil is designed mainly for providing high resolution images for the carotid artery imaging around the carotid bifurcation. Another disadvantage of this design is the combination of the two small 10 cm loop coils into one channel. Combining the two coils into one channel will lower the SNR and resolution of the images compared to using the two coils as two individual channels.
For the three-element mutually decoupled RF coil system (R. Srinivasan, U.S. Pat. No. 6,150,816), the large coil needs to have a size same as the combination of the two small coils. In other words, the large coil cannot be made bigger than the combination of the two small coils. This indicates that the three-element RF coil system can either be used as a high resolution coil by making all the three coils small or be used for covering a large FOV by making all the three coils large. Therefore, the three-element RF coil system cannot be a high resolution and large FOV coil system at the same time. Furthermore, it is not easy to isolate all the three coil elements simultaneously and a complicated iterative testing process is needed for achieving that.
A MRI array coil for imaging a patient having a head, and a torso, includes a base; a left handle extending from the base; a left head coil array attached to the left handle for proximity to the head; a right handle extending from the base; and a right head coil array mounted on the right handle for proximity to the head.